This invention relates to scintillator structures and methods for manufacturing such structures. More particularly, this invention relates to a method of enhancing the escape of visible wavelength radiation from the scintillator structure by matching the index of refraction of the phosphor particles embedded therein with the index of refraction of the optically transparent matrix in which the phosphor is embedded.
In general, a scintillator is a material which emits electromagnetic radiation in the visible spectrum when stimulated by high energy electromagnetic photons such as those in the x-ray or gamma-ray regions of the spectrum, hereinafter referred to as supra-optical frequencies. Thus, these materials are excellent choices for use as detectors in industrial or medical x-ray or gamma-ray equipment. In most typical applications, the light output from scintillator materials is made to impinge upon photoelectrically responsive materials in order to produce an electrical output signal which is in direct relation to the intensity of the initial x-ray or gamma-ray bombardment.
Scintillator materials comprise a major portion of those devices used to detect the presence and intensity of incident high energy photons. The other commonly used detector is the high pressure noble gas ionization device. This other form of high energy photon detector typically contains a gas, such as xenon, at a high pressure (density), which ionizes to a certain extent when subjected to high energy x-ray or gamma-ray radiation. This ionization causes a certain amount of current flow between the cathode and the anode of these detectors which are kept at a relatively high and opposite polarity from one another. The current that flows is sensed by a current sensing circuit whose output is reflective of the intensity of the high energy radiation. Since the high pressure noble gas detector operates on an ionization principle, after the termination of the irradiating energy, there still persists the possibility that a given ionization path remains open through which an undesirable leakage current may pass. Hence, these detectors are peculiarly sensitive to a form of "afterglow" or persistence similar to that found in certain scintillating phosphors. This persistence results in the blurring in the time dimension of the information contained in the irradiating signal.
In general, it is desirable that the amount of light (visible or near visible wavelength) output from these scintillators be as large as possible for a given amount of x-ray or gamma-ray bombardment. This is particularly true in the medical tomography area where it is desired that the energy intensity of the x-ray be as small as possible to minimize the danger to the patient. For this reason the phosphor scintillator should have a good luminescent efficiency.
Another important property that scintillator materials should possess is that of a short afterglow or persistence. This means that there should be a relatively short period of time between the termination of the high energy radiating excitation and the cessation of light output from the scintillator. If this is not the case, there is resultant blurring, in time, of the information-bearing signal. Furthermore, if a rapid scanning is desired, as it is in certain computerized tomographic applications, the presence of the afterglow tends to severely limit the scan rate, thereby rendering difficult the viewing of moving bodily organs, such as the heart or lungs.
A scintillator body or substance, in order to be effective, must be a good converter of high energy radiation (that is, x-rays and gamma-rays). Typically, present scintillator bodies consists of a phosphor in a powder or crystalline form. In this form, the useful light that is produced upon high energy excitation is limited to that which is generated in the surface regions of the body and that which can escape the interior of the scintillator body. This escape is difficult due to multiple internal reflections, each such reflection further attentuating the amount of light externally available by allowing considerably more traversal of phosphor than desired. Thus, it is necessary that not only the phosphors themselves have a good luminescent efficiency but it is also necessary that the light output be available for detection.
In the copending application of Dominic A. Cusano and Jerome S. Prener, Ser. No. 853,086, now U.S. Pat. No. 4,230,510, assigned to the same assignee as this invention, there is described distributed phosphor scintillator structures in which the phosphor is either embedded in an optically transparent matrix or in which the phosphor occurs in a layered structure with alternating layers of phosphor and optically transparent laminate material. This copending application is incorporated by reference herein. In this prior copending application there is still the problem that light rays generated within the scintillator body are refracted and reflected amongst the embedded phosphor particles as a result of the fact that there is a difference in the index of refraction between the phosphor particles and the matrix medium in which they are embedded. This mismatch results in a certain loss of efficiency as measured by light energy escaping the scintillator body.
The term "optical transparency" as used above and hereafter refers to the transparency of the scintillator body or material at or near the wavelength of light emitted by the phosphor or by a single or final wavelength conversion material in the embodiment wherein more than one wavelength conversion material is added. It is to be further noted that the index of refraction of light transmissive materials is in general dependent upon the wavelength of the transmitted light. Thus, the mismatch of indices of refraction mentioned above is a mismatch which is dependent upon the wavelength of light under consideration.
In particular, in the medical tomography area, where the intensity of x-radiation is modulated by the body through which it passes, and which modulated radiation is then converted into electrical signals, it is important to have x-ray detection devices which have as good overall energy conversion efficiency as possible. For devices with low efficiency, a higher flux of x-ray radiation must be applied to produce the same light and electrical output from the overall scintillation detector system. In the context of medical tomography, this means that such systems have a low signal-to-noise ratio.